Biosensors are devices that use biological (cells, enzymes, tissues etc.) to convert a chemical concentration in a matrix into a detectable signal (electrical, acoustic, optical, thermal etc.). There are many types of biosensors used for a wide variety of analytes. Electroenzymatic biosensors which use enzymes to convert a concentration to an electrical signal have been the most studied types of biosensors. For a review of some of the operating principles of biosensors, see P. Bergveld, and D. Thevenot, Advances in Biosensors, Supplement 1, p. 31-94, A. P. F. Turner, editor.
The prototype biosensor is the amperometric glucose sensor. There are several reasons for the wide ranging interest in glucose sensors. The scientific interest is driven by the availability of a very robust enzyme, glucose oxidase, which is used to monitor glucose, as well as the desire to develop model sensors for a wide variety of analytes. The commercial interest is driven by the need for glucose monitoring of patients with diabetes mellitus as well as the development of sensors that can be used to monitor fermentation reactions in the biotechnology arena. A working glucose sensor is also the most difficult component in development of a closed loop artificial pancreas with an implanted insulin pump.
Any amperometric glucose sensor or any oxido-reductase enzyme that uses O.sub.2 as a co-substrate that is designed for subcutaneous or intravenous use requires an outer membrane and an anti-interference membrane because of the fundamental nature of the sensor and the environment in which the measurement is made.
A glucose sensor works according to the following chemical reaction (Equation 1): ##STR1##
In this reaction, glucose reacts with oxygen in the presence of glucose oxidase (GOX) to form gluconolactone and hydrogen peroxide. The gluconolactone reacts with water to open the lactone ring and produce produce gluconic acid. The H.sub.2 O.sub.2 reacts electrochemically as shown below (Equation 2): EQU H.sub.2 O.sub.2 .fwdarw.O.sub.2 +2e.sup.- 2H.sup.+.
The current measured by the sensor/potentiostat (+0.5 to +0.7 V oxidation at Pt black electrode) is due to the two electrons generated by the oxidation of the H.sub.2 O.sub.2. Alternatively, one can measure the decrease in the oxygen by amperometric measurement (-0.5 to -1 V reduction at Pt black electrode).
The stoichiometry of Equation 1 clearly demonstrates some of the problems with an implantable glucose sensor. If there is excess oxygen for Equation 1, then the H.sub.2 O.sub.2 is stoichiometrically related to the amount of glucose that reacts at the enzyme. In this case, the ultimate current is also proportional to the amount of glucose that reacts with the enzyme. If there is insufficient oxygen for all of the glucose to react with the enzyme, then the current will be proportional to the oxygen concentration, not the glucose concentration. For the sensor to be a true glucose sensor, glucose must be the limiting reagent, i.e. the O.sub.2 concentration must be in excess for all potential glucose concentrations. This means that a way must be devised to either increase the O.sub.2 in the GOX membrane, decrease the glucose concentration, or devise a sensor that does not use O.sub.2.
The basic problem in the use of a biosensor in the body is that the ratio of glucose to O.sub.2 is opposite to what is desired for optimal operation of the biosensor. The glucose concentration in the body of a diabetic patient can vary from 2 to 30 mM (millimoles per liter or 36 to 540 mg/dl), whereas the typical oxygen concentration in the tissue is 0.02 to 0.2 mM, U. Fischer, A. Hidde, H. vonWoedtke, K. Rebrin, and P. Abel, Biomed. Biochim. Acta., 1989, Vol. 48, pp. 965-971. This ratio in the body means that the sensor would be running in the Michaelis Menten limited regime and would be very insensitive to small changes in the glucose concentration. This problem has been called the "oxygen deficit problem".
Several approaches to solving the deficit problem have been attempted in the past. The simplest approach is to make a membrane that is fully O.sub.2 permeable, with no glucose permeability and mechanically perforate it with a small hole that allows glucose to pass. Here the differential permeability is defined by the ratio of the small hole area to the total membrane area. Two significant problems with this method are first that reproducibly making small holes is difficult and second and more serious, the O.sub.2 permeability is a strong function of the thickness of the membrane and thickness is difficult to control in mass production. Microporous membranes (U.S. Pat. No. 4,759,828 to Young et al.) have also been tried with limited success. Another problem with both the hole in the membrane approach and the microporous membrane approach is that the sensor electrodes and the enzyme layer are exposed to body fluids. Body fluids contain proteins that coat the electrodes leading to decreased sensitivity of the sensor and enzymes (proteases) that can digest or degrade the sensor active enzyme.
One approach to the oxygen deficit problem is described by Gough (U.S. Pat. No. 4,484,987).
The approach uses a combination membrane with discrete domains of a hydrophilic material embedded in a hydrophobic membrane. In this case, the membrane is not homogenous and manufacturing reproducibility is difficult. Physical properties of the membrane are also compromised. In a similar manner, Gough (U.S. Pat. No. 4,890,620) describes a "two dimensional" system where glucose diffusion is limited to one dimension while the oxygen diffusion is from both dimensions. This sensor is extremely complicated and manufacturing on a large scale is expected to be difficult.
Several other groups, G. W. Shaw, D. J. Claremont and J. C. Pickup, Biosensors and Bioelectronics, 1991, Vol. 6, pp. 401-406; D. S. Bindra, Y. Zhang, G. S. Wilson, R. Sternberg, D. R. Thevenot, D. Moatti and G. Reach, Analytical Chemistry, 1991, Vol. 63, p. 1692; and M. Shichiri, Y. Yamasaki, K. Nao, M. Sekiya and N. Ueda, Horm. Metab. Res., Suppl. Ser., 1988, Vol. 20, p. 17, have used a homogenous membrane of a relatively hydrophobic polyurethane and reported good results. In classical diffusion experiments with these membranes, however, the glucose diffusion is extremely small. It is believed that the ability of these polyurethane layers to allow glucose diffusion is due to micro cracks or micro holes in these materials when applied as membranes.
In order to circumvent the oxygen deficit problem with a homogenous membrane, Allen et al. developed two homogenous membranes with both hydrophilic and hydrophobic regions. In U.S. Pat. No. 5,284,140, they describe an acrylic system and in U.S. Pat. No. 5,322,063 they describe a polyurethane system. Both of the membranes have hydrophilic and hydrophobic moieties in the molecule leading to limited control of oxygen and glucose permeabilities.
The key to stable, high sensitivity enzyme biosensors is that the sensor output must be limited only by the analyte of interest, not by any co-substrates or kinetically controlled parameters such as diffusion. In order to maximize the output current (equation 2) of the biosensor, oxygen diffusion should be as large as possible while maintaining oxygen excess at the reaction surface. Since the normal concentration of O.sub.2 in the subcutaneous tissue is quite low, maximization of the O.sub.2 diffusion coefficient is desirable.
The membrane systems described in the literature as cited above attempt only to circumvent the oxygen deficit problem by reducing the amount of glucose diffusion to the working electrode of the biosensor. The magnitude of the signal from a typical biosensor of the appropriate size for either subcutaneous or intravenous implantation is typically 1 to 10 nA at physiological glucose and O.sub.2 concentrations. This level of current requires sophisticated electronics for measurement. Increasing the oxygen transport and concomitant glucose transport will increase the signal and (to a limited extent) reduce the complexity of the controlling and recording electronics. It is obvious however that a membrane that simultaneously increased the oxygen and limited the glucose would lead to both better performance and increased signal.
Accordingly, there has been a need for a polymer useful in an outer polymeric membrane of a biosensor. There is a need for the membrane to have physical stability and strength, adhesion to the substrate, processibility (ability to be synthesized/manufactured in reasonable quantities and at reasonable prices), biocompatibility, ability to be cut by laser ablation (or some other large scale processing method), and compatibility with the enzyme as deposited on the sensor. The present invention fulfills these needs and provides other related advantages.